The present invention relates to terbium or lutetium containing scintillators that are activated with a rare earth metal ion, such as cerium, and treated by annealing at high temperatures in a defined oxygen atmosphere, such that the annealed scintillator has an increased resistance to high-energy radiation damage as compared to a same scintillator not treated by the methods of the present invention.
Solid state scintillator materials have long been used as radiation detectors to detect penetrating radiation in such applications as x ray counters and image intensifiers. The scintillator materials emit visible or near visible radiation when stimulated by x rays or other high energy electromagnetic photons. In typical medical or industrial applications, the optical output from the scintillator is directed to a photoelectrically responsive device to produce electrical output signals, where the amplitude of the signals is proportional to the deposited energy. The electrical signals can than be digitized by a computer for display on a screen or other permanent medium. Such detectors play an important role in computerized tomography (CT) scanners, digital radiography (DR), and other x ray, gamma medical applications, it is especially desirable that the scintillator efficiently absorb nearly all the x rays that pass through a patient, so that the detector utilizes a maximal amount of the high energy administered, and the patient is not subject to a higher radiation dose than necessary.
Among the preferred scintillator compositions in the present generation of CT scanners are ceramic scintillators that employ at least one of the oxides of lutetium, yttrium, and gadolinium as matrix materials. These are described in detail, for example, in U.S. Pat. Nos. 4,421,671, 4,473,513, 4,525,628, and 4,783,596. These scintillators typically comprise a major proportion of yttria (Y2O3), up to about 50 mole percent gadolinia (Gd2O3), and a minor activating proportion (typically about 0.02–12, preferably about 1–6 and most preferably about 3 mole percent) of a rare earth activator oxide. Suitable activator oxides, as described in the aforementioned patents, include the oxides of europium, neodymium, ytterbium, dysprosium, terbium, and praseodymium. Europium-activated scintillators are often preferred in commercial X ray detectors because of their high luminescent efficiency, low afterglow level, and other favorable characteristics.
Improved scintillator materials for the continually evolving medical applications and technologies that employ radiation-based imaging techniques could be provided. To meet the requirements of typical medical radiographic applications, the scintillator must be an efficient converter of x ray radiation (or other high-energy radiation) into optical radiation for the regions of the electromagnetic spectrum detected by photodetection means. Also, the scintillator should transmit optical radiation efficiently, to avoid trapping of the signal generated within the scintillator body. The scintillator should also be characterized by high x ray stopping power, low hysteresis, spectral linearity, and short afterglow.
One important property of CT systems is scan time, which is the time required for a CT system to scan and acquire an image of a slice of the object under observation. Scan times are related to the primary decay time of the scintillator roughly by a factor of 1000. For example, a scan time of 1 second will typically require a scintillator having a decay time of 1 millisecond or less. Thus, shorter CT scan times require shorter scintillator decay times. The present generation of CT systems have scan times on the order of 1 second, and generally are not lower than about 0.4 second. Still shorter scan times are desired. Decreasing scan time increases the number of patients that can be seen, as well as the number of scans taken in a single measurement, as each measurement requires a patent to “hold their breath” during the measurement period. Shorter scan times also reduce image blurring due to the motion of internal organs or motion that occurs when taking scans of non-cooperating patients, such as small children.
Another consideration for scintillators is to reduce damage that occurs to the scintillator upon repeated exposure of the scintillator to high energy radiation. Radiographic equipment employing solid state scintillator materials for the conversion of high energy radiation to an optical image may experience changes in efficiency after exposure of the scintillator to high dosages of radiation. For example, radiation damage for bismuth germanate single crystal scintillators may be as high as 11% after a thirty minute exposure to ultraviolet radiation from a mercury lamp. Similar results are reported for higher energy gamma radiation. Furthermore, the variation in radiation damage from crystal to crystal of bismuth germanate scintillators is high, approximating a factor of at least 30. A similar change in efficiency can be found when polycrystalline type ceramic scintillators are exposed to high energy radiation dosages.
Radiation damage in scintillators is characterized by a change in light output and/or a darkening in color of the scintillator body with prolonged exposure to radiation. Radiation damage can lead to “ghost images” from prior scans which thereby reduce image resolution. The change in light output that occurs upon radiation damage is often found to be variable in magnitude from batch-to-batch of the same scintillator, making it difficult to predict how any individual scintillator will change over time and thus, making it difficult to implement quantitative correction measures. For example, yttria-gadolinia ceramic scintillators activated with europium exhibit a reduction in light output of 4 to 33%, depending upon the scintillator batch, for 450 roentgens of 140 kVP x rays. This amount of variation in light output which can occur as a result of x ray damage is undesirable in a quantitative x ray detector.